Magnetic resonance imaging method with reduced acoustic noise

ABSTRACT

A magnetic resonance imaging method comprises application of a pulse sequence which includes one or more pulses. The pulse sequence having an intrinsic scan time based on a full sampling rate in k-space for a predetermined full ‘field-of-view’ and a reference temporal pulse shape of the magnetic gradient pulses. A series of magnetic resonance signals is acquired by means of a receiver antennae system having a spatial sensitivity profile. Undersampled signal acquisition is applied to acquire undersampled magnetic resonance signals at a predetermined reduced sampling rate in k-space, the sampling rate being reduced by a reduction factor relative the full sampling rate. The pulse sequence being is during an actual signal scan time is applied. The actual signal scan time being larger than the intrinsic signal scan time times the reduction factor. The undersampling allows a smaller acquisition rate of the magnetic resonance signals and smaller slew rates and amplitudes of the magnetic gradient pulses and lower peak RF-fields of the refocusing pulses. Hence, lower acoustic noise and lower specific absorption rate are achieved.

[0001] The invention relates to a magnetic resonance imaging methodcomprising application of a pulse sequence which includes one or morepules.

[0002] Such a magnetic resonance imaging method is known from the U.S.Pat. No. 4,680,545.

[0003] In general, magnetic resonance imaging methods involve pulsesequences which include radio-frequent (RF) pulses and magnetic gradientpulses. RF-pulses are for example used as excitation pulses and asrefocusing pulses. Magnetic gradient pulses are notably used for spatialencoding of the magnetic resonance signals. Both RF-pulses and magneticgradient pulses can be applied to manipulate the phase of the magneticresonance signals.

[0004] The switching on and off of the magnetic gradient pulses causesacoustic vibrations of gradient coils by which the magnetic gradientsare applied. If no steps are taken, the acoustic vibrations of thegradient coils may cause annoying sounds such as noise, clicks andhumming.

[0005] The known magnetic resonance imaging method employs magneticgradient pulses with a reduced slew rate and sinusoidally shaped pulseedges. Although the known magnetic resonance imaging method reduces theacoustic noise due to the gradient pulses, the reduced slew rate andsinusoidally shaped pulse edges cause to increase the scan time, i.e.the time required to acquire all magnetic resonance signals needed forreconstruction of the magnetic resonance image. Consequently, the knownmagnetic resonance imaging method is not suitable to produce successivemagnetic resonance images at a high rate.

[0006] An object of the invention is to provide a magnetic resonanceimaging method which does not produce annoyingly loud noise and in whichof the scan time is reduced.

[0007] This object is achieved by a magnetic resonance imaging methodaccording to the invention involving acquisition of a series of magneticresonance signals by means of a receiver antennae system having aspatial sensitivity profile, wherein

[0008] undersampled signal acquisition is applied to acquireundersampled magnetic resonance signals at a predetermined reducedsampling rate in k-space, the sampling rate being reduced by a reductionfactor relative the full sampling rate

[0009] the pulse sequence being applied during an actual signal scantime, the actual signal scan time being larger than the intrinsic signalscan time times the reduction factor.

[0010] In order to allow comparison, a reference temporal pulse shape ofthe pulses is introduced. Preferably a commonly used temporal pulseshape, such as a trapezoidal temporal pulse shape may be used as thereference temporal pulse shape for the magnetic gradient pulses, but anypredetermined temporal pulse shape can be employed as the referencetemporal pulse shape. For the RF-pulses, notably the RF-excitationpulses and the refocusing RF-pulses, a reference pulses shape isrepresented for example by its temporal pulse duration, peak RF-fieldamplitude and the flip angle over which spins are flipped by theRF-pulse at issue. It is noted that the flip angle depends on temporalpulse duration and peak RF-amplitude.

[0011] Further, the reference temporal pulse shape may involve differentor the same pulse shapes for respective individual magnetic gradientpulses and/or RF-pulses in the pulse sequence.

[0012] According to the invention, the magnetic resonance signals areacquired at a sampling rate in k-space that is less than required forthe predetermined full ‘field-of-view’. As a consequence of the reducedsampling rate, these undersampled magnetic resonance signals only have apartial spatial encoding in that individual undersampled magneticresonance signals include contributions from different spatial positionswithin the predetermined full ‘field-of-view’. Signal contributionsrelating to separate spatial positions are computed from theundersampled magnetic resonance signals on the basis of the spatialsensitivity profile of the receiver antennae system. In other words, thespatial sensitivity profile of the receiver antennae system provides theadditional spatial encoding that is not present in the undersampledmagnetic resonance signals. A magnetic resonance image is reconstructedfrom the undersampled magnetic resonance signals and with the use of thespatial sensitivity profile of the receiver antennae system.

[0013] The undersampling allows the acquisition of less undersampledmagnetic resonance signals than the full sampling would require.According to the magnetic resonance imaging method according to theinvention, the scan time of the pulse sequence is less reduced, or notreduced at all, than the reduction of the sampling rate in k-space wouldallow for acquisition of the same number of magnetic resonance signals.Hence, the invention allows the magnetic resonance signals to beacquired at a lower temporal rate and consequently the temporalvariations of the magnetic gradient pulses in the pulse sequence aremore slowly than the reference temporal pulse shape. In order to make afair comparison, the intrinsic signal scan time is defined for a pulsesequence that only differs from the actually employed pulse sequence asto the temporal reference pulse shape of the magnetic gradient pulsesand of the RF-pulses. That is, the intrinsic signal scan time iscomputed for a pulse sequence having the same succession of magneticgradient pulses and RF-pulses as the actual pulse sequence, but whereinthe magnetic gradient pulses have the reference temporal pulse shape(s).Thus, less acoustic noise is generated, while the scan time is equal toor still less than the intrinsic scan time that is involved if themagnetic gradient pulses with the reference pulse shaped are employed.Notably, the acoustic noise is less as the actual signal scan time isreduced less than what the reduction of the sampling rate in k-spaceallows. In particular, the invention allows to obtain an optimumcompromise between reduction of the actual signal scan time andreduction of acoustic noise for different circumstances. For example,when the gradient slopes rise to 10 mTm⁻¹ in 1-3 ms are employed, ascompared to a gradient slope of 21 mTm⁻¹ in 0.2 ms for the referencepulse shape, the acoustic noise is reduced by about 20-30 dB.

[0014] Moreover, the undersampling allows to use a lower gradientamplitude, using a lower receiver bandwidth owing to the correspondinglylonger signal scan time. This will increase the signal-to-noise ratio ofthe magnetic resonance signals.

[0015] The degree of undersampling is given by the reduction factorwhich is defined as the ratio of the actual sampling rate or density ink-space relative to the full sampling rate. Thus, if e.g. only half ofthe magnetic resonance signals required for full sampling is actuallysampled, the reduction factor is ½. More generally, the reduction factoris a rational number in the interval [0,1].

[0016] In addition of the reduction of the acoustic noise, alsoreduction of the electromagnetic energy deposited by the pulse sequencein the patient's tissue is reduced. This deposition of electromagneticenergy is represented by the quantity usually denoted ‘specificabsorption rate’. The specific absorption rate (or SAR) is defined asthe radio frequency power absorbed per unit of mass of an object. Inparticular, the specific absorption rate is reduced owing to reductionof the slew rate of the magnetic gradient pulses. Also lower peakRF-fields of the RF-pulses reduce the specific absorption rate. Since,the scan time is less reduced than the degree of undersampling of themagnetic resonance signals would allow, the temporal duration of theRF-pulses can be increased so as to achieve the desired flip angle whileemploying a lower peak RF-field without the need the increase the scantime relative to the reference pulse sequence. The reduction of thespecific absorption rate achieved by the magnetic resonance imagingmethod according to the invention causes less inadvertent heating of thepatient's local body temperature. Hence, physiological uncomfortable oreven detrimental effects caused by increase of the patient's bodytemperature are mitigated or even avoided. The lower specific absorptionrate in combination with the possibility to employ relatively short scantime is especially more advantageous as stronger static magnetic fieldsare employed, such as in a field strength region above 1.5 T, especiallyin magnetic fields of about 3 T. In addition, reduction of the slew rateand/or of the gradient amplitude of the magnetic gradient pulses resultsin lowering of peripheral nerve stimulation in the patient to beexamined.

[0017] These and other aspects of the invention will be furtherelaborated with reference to the embodiments defined in the dependentClaims.

[0018] A particular effective reduction of acoustic noise and/orspecific absorption rate is achieved when the actual signal scan time isequal or about equal to the intrinsic signal scan time. In thisimplementation the reduction of the number of magnetic resonance signalsacquired is entirely used to reduce the signal acquisition rate andconsequently to reduce the temporal variations of the magnetic gradientpulses and/or to reduce the number and peak RF-field of the refocusingRF-pulses. Nevertheless, the actual signal scan time does not yet exceedthe intrinsic signal scan time.

[0019] In particular, effective noise reduction is achieved by employingmagnetic gradient pulses having temporal variations that are more slowlythan the reference temporal variations. Although the time involved toapply the magnetic gradient pulses is longer, the time required toacquire the magnetic resonance signals is not increased because lessmagnetic resonance signals are acquired owing to the undersampling.

[0020] For example, noise reduction is achieved by employing a slew rateof the magnetic gradient pulses that is lower than the slew rate in thereference temporal pulse shape. As the slew rate is one of the dominantsources of acoustic vibrations, decreasing the slew rate leads to asubstantial reduction of acoustic noise.

[0021] The number of acquired magnetic resonance (MR) signals for anindividual magnetic resonance image is reduced by employingundersampling of the MR-signals. Such undersampling involves a reductionin k-space of the number of sampled points which can be achieved invarious ways. Many magnetic resonance imaging methods which involveundersampling have been proposed. Often these magnetic resonance imagingmethods are termed ‘parallel imaging methods’, because the undersampledmagnetic resonance signals relate to several lines in k-spacesimultaneously, so that in fact two or more lines in k-space areacquired in parallel. Notably, the MR signals are picked-up throughsignal channels pertaining to several receiver antennae, such asreceiver coils, preferably surface coils. Acquisition through severalsignal channels enables parallel acquisition of signals so as to furtherreduce the signal scan time.

[0022] Owing to the undersampling, sampled data contain contributionsfrom several positions in the object being imaged. The magneticresonance image is reconstructed from the undersampled MR-signals withthe use of a sensitivity profile associated with the signal channels.Notably, the sensitivity profile is for example the spatial sensitivityprofile of the receiver antennae, such as receiver coils. Preferably,surface coils are employed as the receiver antennae. The reconstructedmagnetic resonance image may be considered as being composed of a largenumber of spatial harmonic components which are associated withbrightness/contrast variations at respective wavelengths. The resolutionof the magnetic resonance image is determined by the smallestwavelength, that is by the highest wavenumber (k-value).The largestwavelength, i.e. the smallest wavenumber, involved, is the field-of-view(FOV) of the magnetic resonance image. The resolution is determined bythe ratio of the field-of-view and the number of samples.

[0023] The undersampling may be achieved in that respective receiverantennae acquire MR signals such that their resolution in k-space iscoarser than required for the resolution of the magnetic resonanceimage. The smallest wavenumber sampled, i.e. the minimum step-size ink-space, is increased while the largest wavenumber sampled ismaintained. Hence, the image resolution remains the same when applyingundersampling, while the minimum k-space step increases, i.e. the FOVdecreases. The undersampling may be achieved by reduction of the sampledensity in k-space, for instance by skipping lines in the scanning ofk-space so that lines in k-space are scanned which are more widelyseparated than required for the resolution of the magnetic resonanceimage. The undersampling may be achieved by reducing the field-of-viewwhile maintaining the largest k-value so that the number of sampledpoints is accordingly reduced. Owing to the reduced field-of-viewsampled data contain contributions from several positions in the objectbeing imaged.

[0024] Notably, when receiver coil images are reconstructed fromundersampled MR-signals from respective receiver coils, such receivercoil images contain aliasing artefacts caused by the reducedfield-of-view. From the receiver coil images and the sensitivityprofiles the contributions in individual positions of the receiver coilimages from different positions in the image are disentangled and themagnetic resonance image is reconstructed. This MR-imaging method isknown as such under the acronym SENSE-method. This SENSE-method isdiscussed in more detail in the international application no. WO99/54746-A1.

[0025] Alternatively, the undersampled MR-signals may be combined intocombined MR-signals which provide sampling of k-space corresponding tothe full field-of-view. In particular, according to the so-calledSMASH-method undersampled MR-signals signals approximate low-orderspherical harmonics which are combined according to the sensitivityprofiles. The SMASH-method is known as such from the internationalapplication no. WO 98/21600.

[0026] Undersampling may also be carried-out spatially. In that case thespatial resolution of the MR-signals is less than the resolution of themagnetic resonance image and MR-signals corresponding to a fullresolution of the magnetic resonance image are formed on the basis ofthe sensitivity profile. Spatial undersampling is in particular achievedin that MR-signals in separate signal channels, e.g. from individualreceiver coils, form a combination of contributions from severalportions of the object. Such portions are for example simultaneouslyexcited slices. Often the MR-signals in each signal channel form linearcombinations of contributions from several portions, e.g. slices. Thislinear combination involves the sensitivity profile associated with thesignal channels, i.e. of the receiver coils. Thus, the MR-signals of therespective signal channels and the MR-signals of respective portions(slices) are related by a sensitivity matrix which represents weights ofthe contribution of several portions of the object in the respectivesignal channels due to the sensitivity profile. By inversion of thesensitivity matrix, MR-signals pertaining to respective portions of theobject are derived. In particular MR-signals from respective slices arederived and magnetic resonance images of these slices are reconstructed.

[0027] Advantageously, in the pulse sequence of the magnetic resonanceimaging method according to the invention the pulse shape(s) of themagnetic gradient pulses is adjustable. Thus it is achieved that thesignal scan time can be reduced at the cost of a higher acoustic noiselevel and/or specific absorption rate. In some examinationcircumstances, having a short signal scan time is more worthwhile thanlowering the acoustic noise level and/or specific absorption rate.Particular examples are magnetic resonance imaging methods which involvescanning within a patient's breathhold, magnetic resonance imagingmethods aimed at imaging perfusion in the patient's brain and in generalmagnetic resonance imaging methods driven by temporal physiologicalprocesses in the patient to be examined. The adjustable actual signalscan time renders the magnetic resonance imaging method according to theinvention more flexible than the conventional magnetic resonance imagingmethod.

[0028] These and other aspects of the invention will be elucidated withreference to the embodiments described hereinafter and with reference tothe accompanying drawing wherein

[0029]FIG. 1 shows diagrammatically a magnetic resonance imaging systemin which the invention is used and

[0030] FIGS. 2 to 5 show simple examples of comparison of the actualsignal scan time to the intrinsic scan time.

[0031]FIG. 1 shows diagrammatically a magnetic resonance imaging systemin which the invention is used. The magnetic resonance imaging systemincludes a set of main coils 10 whereby the steady, uniform magneticfield is generated. The main coils are constructed, for example in sucha manner that they enclose a tunnel-shaped examination space. Thepatient to be examined is slid into this tunnel-shaped examinationspace. The magnetic resonance imaging system also includes a number ofgradient coils 11, 12 whereby magnetic fields exhibiting spatialvariations, notably in the form of temporary gradients in individualdirections, are generated so as to be superposed on the uniform magneticfield. The gradient coils 11, 12 are connected to a controllable powersupply unit 21. the gradient coils 11, 12 are energized by applicationof an electric current by means of the power supply unit 21. Thestrength, direction and duration of the gradients are controlled bycontrol of the power supply unit. The magnetic resonance imaging systemalso includes transmission and receiving coils 13, 16 for generating theRF excitation pulses and for picking up the magnetic resonance signals,respectively. The transmission coil 13 is preferably constructed as abody coil 13 whereby (a part of) the object to be examined can beenclosed. The body coil is usually arranged in the magnetic resonanceimaging system in such a manner that the patient 30 to be examined isenclosed by the body coil 13 when he or she is arranged in the magneticresonance imaging system. The body coil 13 acts as a transmissionantenna for the transmission of the RF excitation pulses and RFrefocusing pulses. Preferably, the body coil 13 involves a spatiallyuniform intensity distribution of the transmitted RF pulses (RFS). Thesame coil or antenna is usually used alternately as the transmissioncoil and the receiving coil. Preferably, a so-called synergy coil isemployed as the body coil. Furthermore, the transmission and receivingcoil is usually shaped as a coil, but other geometries where thetransmission and receiving coil acts as a transmission and receivingantenna for RF electromagnetic signals are also feasible. Thetransmission and receiving coil 13 is connected to an electronictransmission and receiving circuit 15.

[0032] It is to be noted that preferably separate receiving coils 16 areemployed. In particular, surface coils 16 can be used as receivingcoils. Such surface coils have a high sensitivity in a comparativelysmall volume. The spatial sensitivity profiles of the surface coils arepreferably calibrated relative to the uniform sensitivity profile of thebody coil.

[0033] The transmission coils, such as the surface coils, are connectedto a demodulator 24 and the received magnetic resonance signals (MS) aredemodulated by means of the demodulator 24. The demodulated magneticresonance signals (DMS) are applied to a reconstruction unit. Thereceiving coil is connected to a preamplifier 23. The preamplifier 23amplifies the RF resonance signal (MS) received by the receiving coil 16and the amplified RF resonance signal is applied to a demodulator 24.The demodulator 24 demodulates the amplified RF resonance signal. Thedemodulated resonance signal contains the actual information concerningthe local spin densities in the part of the object to be imaged.Furthermore, the transmission and receiving circuit 15 is connected to amodulator 22. The modulator 22 and the transmission and receivingcircuit 15 activate the transmission coil 13 so as to transmit the RFexcitation and refocusing pulses. The reconstruction unit derives one ormore image signals from the demodulated magnetic resonance signals(DMS), which image signals represent the image information of the imagedpart of the object to be examined. The reconstruction unit 25 inpractice is constructed preferably as a digital image processing unit 25which is programmed so as to derive from the demodulated magneticresonance signals the image signals which represent the imageinformation of the part of the object to be imaged. The signal on theoutput of the reconstruction monitor 26, so that the monitor can displaythe magnetic resonance image. It is alternatively possible to store thesignal from the reconstruction unit 25 in a buffer unit 27 whileawaiting further processing.

[0034] The magnetic resonance imaging system according to the inventionis also provided with a control unit 20, for example in the form of acomputer which includes a (micro)processor. The control unit 20 controlsthe execution of the RF excitations and the application of the temporarygradient fields. In particular, according to the invention, the controlunit is arranged to adjust the pulse shapes of the magnetic gradientpulses, such as the read gradient pulses and the phase-encoding gradientpulses. The control unit is arranged to receive a command from the userto set up the pulse sequence such that acoustic noise reduction isachieved. Upon this command, the control unit sets up the pulse sequencesuch that undersampled scanning of k-space is applied and e.g. the SENSEtechnique is applied for the reconstruction of the magnetic resonanceimage. Further, the pulse shapes are adjusted such that the signal scantime of the pulse sequence is about equal to the intrinsic signal scantime. To this end, the computer program according to the invention isloaded, for example, into the control unit 20 and the reconstructionunit 25.

[0035] FIGS. 2 to 5 show a simple examples of comparison of the actualsignal scan time to the intrinsic scan time. The exemplary pulsesequences shown graphically and by way of example in FIGS. 2 to 5 areTSE sequences. FIG. 2 represents the reference TSE pulse sequence forthis example. Graph indicated ‘G_(read)’ in FIG. 2 indicates the readgradient pulses in the read direction and also the magnetic resonancesignals are represented. Graph G_(slice) indicates the slice magneticselection gradient pulses. G_(enc) indicates the magnetic phase-encodinggradient pulses. Graph RF indicates the RF-pulses, such asradio-frequent excitation and refocusing pulses. The pulse sequencestarts with an excitation RF-pulse (RFer, α₁), e.g. a 90°-pulse toexcite magnetic spins in the object to be examined. The RF-excitationpulse (RFer) acts slice selectively through a magnetic slice selectiongradient pulse (Gs1 r) that is applied simultaneously with theRF-excitation pulse (RFer). Subsequently, successive RF refocusingpulses (Rf1 r, Rf2 r,. . . Rf4 r) (β₁) are applied to refocusing thetransverse magnetization component are applied to generate spin-echomagnetic resonance signals. Owing to these RF refocusing pulses, asuccession of spin echo magnetic resonance signals (MR1 r, MR2 r, MR3 r,MR4 r) occurs. To achieve spatial encoding, phase encoding magneticgradient pulses (PE1 r, PE2 r,. . . PE8 r) and read-gradient pulses (RG1r, RG2 r, RG3 r, RG4 r and RG5 r) are applied as shown in graph‘G_(enc)’ of FIG. 2. The read gradient pulses achieve scanning ofk-space in the k_(x)-direction. Subsequent phase encoding pulses (PE1 r,PE2 r,. . . PE8 r) provide respective shifts in the k_(y)-direction. Thetemporal shape of the various pulse shapes of the read gradient pulsesand phase encoding gradient pulses have the reference temporal pulseshape(s). Notably, the reference temporal pulse shapes have relativelyhigh slew rates, so as to allow a relatively short intrinsic scan time(T_(R0)).

[0036]FIG. 2 shows a schematic representation of the gradient and RFwave-forms forms a typical 4 ETL TSE sequence with a fixed repetitiontime equal to T_(R). The total signal scan time is thereforeT_(scan)=N×T_(R) where N is the number of interleaves required toacquire a complete image matrix. For a 256 encoded image matrix N=64.This signal scan time T_(scan) of the reference pulse sequence is theintrinsic signal scan time for this 4 ETL TSE sequence. Gradient slewrate is set to maximum in this example. $\begin{matrix}{{SAR} \propto \frac{{\sum\limits_{T_{scan}}{{\alpha 1}( {{B1}_{\max},t} )}^{2}} + {\sum\limits_{T_{scan}}{{\beta 1}( {{B1}_{\max},t} )}^{2}}}{T_{scan}}} & \lbrack 1\rbrack\end{matrix}$

[0037] Typically β1=2·α1 where β1 is a 180 degree RF pulse which is afunction of pulse shape, duration, t, and peak RF field amplitude equalto some maximum available value B1_(max) μT. In this example the totalnumber of α1 pulses is 64 and the total number of β1 pulses is 4×64=256for the fully encoded scan. For fixed TR, acoustic noise is proportionalto gradient amplitude, slew rate and the total number of gradientslopes. Assuming all diagrams are drawn on the same scale, then in FIG.2, the gradient amplitudes represent initial amplitudes as required bythe parameters of the TSE sequence. The gradient rise-time isinstantaneous representing maximum slew rate. In general then:

Acoustic_Noise_Level ∝G·S·N_(slopes)   [2]

[0038] Where G=gradient amplitude, S=gradient slew rate andN_(slopes)=No. of gradient slopes. Reducing either G,S,N_(slopes) or allthree simultaneously, can lead to a significant lowering of acousticnoise.

[0039] Graph indicated ‘Gread’ in FIG. 3 shows the pulse shapes of theread gradient pulses in an example of the actual the pulse sequencewhere undersampling is applied. Graph G_(slice) indicates the slicemagnetic selection gradient pulses. Graph G_(enc) indicates the magneticphase-encoding gradient pulses. Graph RF indicates the RF-pulses, suchas radio-frequent excitation and refocusing pulses. The pulse sequenceshown in FIG. 3 has less gradient pulses having lower amplitudes andslew rates relative to the pulse sequence shown in FIG. 2 having theintrinsic signal scan time and the reference pulse shapes. The number ofRF refocusing pulses (Rf1 a,. . . Rf2 a) which are applied to refocusingthe transverse magnetization component is smaller, about half of thenumber of RF refocusing pulses in the reference TSE pulse sequence shownin FIG. 2. The read gradient pulses (RG1 a, RG2 a, RG3 a) and themagnetic slice selection gradient (Gs1 a)in FIG. 3 have a far smallerslew rate relative to the slew rate of the reference pulse shapes shownin FIG. 2. Further, less spin echo magnetic resonance signals (MR1 a,MR2 a) are generated. Also the phase encoding gradient pulses (PE1 a, .. . PE4 a)) having lower slew rates and amplitudes. The repetition timeT_(R) is the same as for the reference pulse sequence of FIG. 2. Becauseless, in this example only half, of the number of magnetic resonancesignals are acquired the actual signal scan time equals the intrinsicsignal scan time (T_(scan)).

[0040]FIG. 3 is similar to FIG. 2, except the gradient slew rate of boththe magnetic read gradient pulses and of the magnetic slice selectiongradient pulses is now reduced. This has the effect of reduced acousticnoise in addition to the reduced SAR. Since the gradient slew rate isreduced it is necessary to space the gradient wave-forms further apart.Since the repetition time T_(R) is fixed and there is spare timeavailable, this is now possible with no penalty on the total signal scantime, which remains the same T_(scan), as in FIG. 2. In this example,SENSE is used to complete the full matrix acquisition. A furtheradvantage of reducing the slew rate is the effect it has in lowering thetime derivative of the magnetic (gradient) field dB/dt which can be acause of peripheral nerve stimulation (PNS) in fast imaging sequencesthat use strong gradients.

[0041] Graph indicated ‘G_(read)’ in FIG. 4 shows the pulse shapes ofthe read gradient pulses in an example of the actual the pulse sequencewhere undersampling is applied. Graph G_(slice) indicates the slicemagnetic selection gradient pulses. Graph G_(enc) indicates the magneticphase-encoding gradient pulses. Graph RF indicates the RF-pulses, suchas radio-frequent excitation and refocusing pulses. The pulse sequenceshown in FIG. 4 has less gradient pulses having lower amplitudes andslew rates relative to the pulse sequence shown in FIG. 2 having theintrinsic signal scan time and the reference pulse shapes. The readgradient pulses (RG1 a, RG2 b, RG3 b) and the magnetic slice selectiongradient (Gs1 a) in FIG. 4 have a far smaller slew rate relative to theslew rate of the reference pulse shapes shown in FIG. 2. Further, theread gradient pulses (RG2 b, RG3 b) have been decreased in amplitude andlengthened in time. The actual signal scan time (T_(scan)) equals theintrinsic signal scan time.

[0042]FIG. 4 shows a sequence similar to that of FIG. 3 except the“G_(read)” gradient wave-forms have now been decreased in amplitude andlengthened in time. The extra time available in the T_(R) is used toaccommodate the resulting increase in echo spacing. The net effect oflowering and lengthening the “G read” wave-forms is threefold.

[0043] 1. The lowering of gradient amplitude reduces further theacoustic noise.

[0044] 2. The lowering of gradient amplitude reduces further thepossibility of PNS.

[0045] 3. The lengthening of the “G read” wave-forms is compensated byusing a lower acquisition bandwidth which results in a higher SNR inmany cases.

[0046] The SENSE technique is again utilized to recover the full 256matrix from the 64×2=128 encoding steps that are actually acquired forreconstruction of the magnetic resonance image.

[0047] The SAR is identical to that of FIG. 3 (lower than the scan usingthe reference sequence), the acoustic noise and PNS potential are evenlower than that of FIG. 3 and the SNR is improved due to the lowerbandwidth now used to sample each echo. Once again, the total scan time,T_(scan), and image resolution/properties are essentially equal to thoseof the scan using the reference pulse sequence of FIG. 2.

[0048] Graph indicated ‘G_(read)’ in FIG. 5 shows the pulse shapes ofthe read gradient pulses in an example of the actual the pulse sequencewhere undersampling is applied. Graph G_(slice) indicates the slicemagnetic selection gradient pulses. Graph G_(enc) indicates the magneticphase-encoding gradient pulses. Graph RF indicates the RF-pulses, suchas radio-frequent excitation and refocusing pulses. The pulse sequenceshown in FIG. 5 has less gradient pulses having lower amplitudes andslew rates relative to the pulse sequence shown in FIG. 2 having theintrinsic signal scan time and the reference pulse shapes. The readgradient pulses (RG1 a, RG2 b, RG3 b) and the magnetic slice selectiongradient (Gs1 b) in FIG. 5 have a far smaller slew rate relative to theslew rate of the reference pulse shapes shown in FIG. 2. Further, thepeak RF-fields of the RF-excitation pulse (Rfeb) and of refocusingRF-pulses (Rf1 b, RF2 b) are lower as compared to the reference pulsesequence. Also, as in the example in FIG. 4, the read gradient pulses(RG2 b,RG3 b) have been decreased in amplitude and lengthened in timeThe actual signal scan time is about equal to the intrinsic signal scantime.

[0049]FIG. 5 is similar to FIG. 4. In this case, RF pulses that use alower B1 are employed, (represented by α₂ and β₂).

[0050] The use of a lower B1 further reduces the SAR of the sequence. Inaddition, it requires to also reduce the RF pulse bandwidth whichresults in longer duration RF pulses. As a result of this reducedbandwidth, the “G slice ” gradient is also reduced in amplitude. Thisreduction in amplitude leads to a further reduction in acoustic noiselevel and PNS potential.

[0051] Since the RF pulses are now longer in duration they are spacedfurther apart. Utilizing the remaining time that is left within the TR,it is possible to space out the RF pulses further and additionallylength and lower the “G_(read)” gradient wave-forms. This secondaryeffect provides both a further reduction in acoustic noise levels and afurther increase in SNR.

[0052] As in all previous embodiments, the total acquisition timeremains the same and the SENSE method is utilizing to generate the datafor the full 256 matrix.

[0053] The PNS potential of this implementation is yet lower thanprevious examples.

1. A magnetic resonance imaging method comprising application of a pulsesequence which includes one or more pulses the pulse sequence having anintrinsic scan time based on a full sampling rate in k-space for apredetermined full ‘field-of-view ’ and a reference temporal pulse shapeof the magnetic gradient pulses acquisition of a series of magneticresonance signals by means of a receiver antennae system having aspatial sensitivity profile, wherein undersampled signal acquisition isapplied to acquire undersampled magnetic resonance signals at apredetermined reduced sampling rate in k-space, the sampling rate beingreduced by a reduction factor relative the full sampling rate the pulsesequence being applied during an actual signal scan time, the actualsignal scan time being larger than the intrinsic signal scan time timesthe reduction factor.
 2. A magnetic resonance imaging method as claimedin claim 1, wherein the actual signal scan time is substantially equalto the intrinsic signal scan time.
 3. A magnetic resonance imagingmethod as claimed in claim 1, wherein the pulses are applied with apulse shape having temporal variations that are more slowly than thetemporal variations involved in the reference temporal pulse shape.
 4. Amagnetic resonance imaging method as claimed in claim 4, wherein thepulse sequence includes magnetic gradient pulses and the magneticgradient pulses are applied with a slew rate that is less than the slewrate involved with the reference pulse shape.
 5. A magnetic resonanceimaging method as claimed in claim 4, wherein the pulse sequenceincludes radiofrequent (RF) pulses and the radiofrequent (RF) pulses areapplied with an RF-magnetic field component (B₁) that is less than theRF-magnetic field component (B₁) involved with the reference pulseshape.
 6. A magnetic resonance imaging method as claimed in claim 1,wherein the pulse shape of the pulses is adjustable so as to adapt theactual signal scan time of the pulse sequence in the range between theintrinsic signal scan time times the reduction factor and the intrinsicsignal scan time.